This invention generally relates to instrumentation for sensing various blood analytes, and more particularly towards optical sensors for detecting blood components which are substrates for oxidative enzymes.
Various sensors have been described in the art for measuring the presence of a substrate which may be enzymatically oxidized. Permanent or multiple use sensors have been made, for example, as in U.S. Pat. No. 5,494,562, disclosing an electrochemical glucose sensor wherein the enzyme glucose oxidase is co-immobilized in a platinized carbon paste in contact with an electrode. The paste is further covered with a semipermeable porous polymer to permit the movement of hydrated glucose and oxygen gas to the paste, where the active enzyme converts these substrates to peroxide and gluconate. The polarized electrode then delivers a current and the reoxidation of peroxide to oxygen may be measured amperometrically. Examples of other amperometric sensor technologies have been described in U.S. Pat. Nos. 4,680,268, 5,352,348, and 5,298,144. These sensors are disadvantageous for use in disposable applications, because they were designed to function in a multiuse format with an elaborate instrument system, and they employ expensive materials and methods of manufacture as well as requiring wet up and calibration before use.
Another example of permanent or multiple use sensors, using fiber optic glucose sensors, is disclosed in U.S. Pat. No. 4,974,929, describing a chamber filled with an oxygen quenchable dye at the end of a fiber surrounded by a double jacketed wall. Sandwiched between the wall jackets is a crosslinked polymer layer containing the enzyme glucose oxidase. The glucose and oxygen consumption at the cylinder wall lowers oxygen diffusion into the cylinder space. This sensor requires a fair degree of manual assembly and is also not suitable for inexpensive, routine single use measurement applications. Each sensor must also be individually calibrated before use, and employs a large cylindrical volume for dye equilibration and accordingly suffers in response speed due to the requirement for bulk equilibration effects.
Disposable or "one-shot" sensors, on the other hand, have been disclosed also, such as the colorimetric sensor set forth in U.S. Pat. No. 5,208,147. Disposable dry reagent strips, e.g., U.S. Pat. Nos. 3,992,158, 4,689,309, and 5,520,883, were developed specifically for single use applications. In operation, the sample hydrates the test strip and reagents are consumed in the development of a colorimetric change based on peroxide chemistries. They can be stored dry, are ready to use on demand, and find use in "wet" blood or serum chemistry, where the strips become saturated during use. However, the hydration and depletion of reactive chemical reagents in these strips effectively prevents their re-use. Handling and biohazard disposal of the many individually used test elements that would routinely be generated is a drawback.
In the development of optical sensors which are inexpensive to manufacture and may be used in a rapid point of care setting, optical coatings and membranes containing luminescent dyes on light-transmissive substrates for the detection and measurement of O.sub.2 in blood, have been developed and are incorporated by reference as U.S. application Ser. No. 08/617,714. An advantage of these sensors is that it is possible to bring the sensor into intimate contact with a blood sample while transmitting excitation light through the transparent substrate from the "back" side of the sensor, to enable subsequent detection of the luminescent radiation emitted from the luminescent dye from the same side of the sensor. Such coatings or membranes are quite thin, typically 1-5 .mu.m for pH and oxygen sensors, and show an extremely rapid response.
However, the application of such thin optical coating-based technologies to the area of glucose determination is not straightforward. Firstly, measurements of oxygen uptake in blood are directly affected by the oxygen buffering capacity of a blood sample itself. If blood is introduced into a measurement system containing an optical sensor, red blood cells (RBCs) will be in contact with the sensor. Since the principal of operation of optical glucose sensors is based on oxidization of available glucose in the blood sample by immobilized glucose oxidase in the sensor (i.e., the detection of oxygen is directly correlatable to the glucose concentration), the depletion of oxygen proximate to the membrane will be high. As such, red blood cells (RBCs) in contact with the sensor will sense the low oxygen concentration at the membrane surface and respond by releasing bound oxygen, thus acting as an oxygen buffer and skewing the results obtained.
Secondly, a large fraction of the excitation light (in some cases as much as 90-95%) as well as the luminescent energy emitted from the sensing layer from these thin membranes passes into the sample (blood) to be absorbed, scattered, or reflected back into the sensing layer and into the detection optics. These effects can, together, produce a four-fold change in luminescence signal level between a perfectly absorbing and perfectly reflecting sample, thus presenting a significant source of uncertainty in the use of luminescent optical sensors, particularly if used in the simplest amplitude readout mode.
Accordingly, it has been desired to obtain optical sensors for measuring oxidizable analytes which are inexpensive enough to design and manufacture to make them disposable, avoid the aforementioned drawbacks, and also allow multiple uses if desired.